Locomotive implantable devices have numerous applications including sensing, imaging, minimally invasive surgery, and research. Many techniques have been used to generate motion, including mechanical solutions and passive magnetic solutions. Power sources dominate the size of existing active implant technologies, and this size constraint (typically in the cm-range) limits the potential for propulsion. Additionally, mechanical propulsion is inherently inefficient at the scale of interest.
Passive locomotion schemes have circumvented the power and efficiency issues, but require large field gradients and usually cannot generate vertical motion. In a passive magnetic propulsion technique, a force is exerted on a small ferromagnetic material with magnetic field gradients. The passive propulsion method typically employs MRI-like systems because the gradient fields must be large and precisely controlled. The gradient must be in the direction of movement, and even MRI systems cannot overcome the force of gravity for devices smaller than roughly 1 mm. The force scales poorly as the size is reduced because it is proportional to the volume of the object. From a practical perspective, generating large field gradients is complicated, and current technology is inadequate.
In addition to the passive method, it is also possible to use mechanical propulsion with active power. Mechanical propulsion is accomplished with a wide variety of techniques. A few possible methods include flagella/motors, pumps, and acoustic streaming. These designs typically suffer from low conversion efficiency from input power to thrust, especially as the Reynolds number decreases. There are losses associated with the conversion from electrical power to mechanical motion, and more loss associated with the conversion from mechanical motion to forward thrust. As a result of the low efficiency, a fairly substantial amount of power is required, and the power source dominates the size making it difficult to miniaturize.
Moreover, Implantable medical devices (IMDs), such as locomotive implantable devices, are a rapidly growing area of technology. In-vivo monitoring and treatment of key biological parameters can greatly assist in managing health and preventing disease. IMDs are complete systems often incorporating signal transducers, wireless data transceivers and signal processing circuits. Power consumption in these devices requires batteries, which must be replaced periodically, or inductive power coupling antennae, both of which dominate device volume, increasing patient discomfort and severely restricting the range of viable applications.
Previous inductive powering links for IMDs operate in the low MHz requiring loop antenna diameters of a few cm and near-perfect transmitter and receiver alignment to deliver sufficient power. This choice of frequency is usually explained by saying that tissue losses become too large at higher frequencies and referring to a qualitative analysis. For these low MHz inductively coupled links the range is much less than a wavelength and thus the links satisfy the near field approximation to Maxwell's equations. Therefore resonant tuning techniques can be used to achieve the maximum energy transfer from the source to the load circuits for these links. Inductive coupling antennae of this size are viable for retinal implants where there is an existing cavity in the eye-socket but are much too large for many other IMDs such as implantable glucose sensors.
The physics behind wireless powering is described first. A time-varying current is set up on the transmit antenna. This gives rise to a time-varying magnetic field. The time-varying magnetic field, in turn, gives rise to an electric field. The electric field induces a current on the receive antenna. Then, this induced current on the receive antenna intercepts the incident electric field and/or magnetic field from the transmit antenna, and generates power at the receiver. Prior devices for wireless transmission of power to medical implants mainly operate based on inductive coupling over the near field in conjunction with a few based on electromagnetic radiation over the far field.
Devices based on inductive coupling operate at very low frequency, 10 kHz to 1 MHz. A wavelength is long relative to the size of the transmit and receive antennas. They are usually a few cm in diameter. Most energy stored in the field generated by the transmit antenna is reactive, that is, the energy will go back to the transmitter if there is no receiver to intercept the field. The separation between transmit and receive antennas is very small, usually a few mm. The low frequency and the short separation mean that there is apparently no phase change between the field at the transmitter and the incident field at the receiver. The increase in the transmit power due to the presence of the receiver mostly delivers to the receiver, like a transformer. Prior devices are therefore designed using the transformer model where various tuning techniques are proposed.
To deliver sufficient power to the implant using inductive coupling based devices, the receive antenna attached to the implant is of a few cm in diameter which is too large. It is required to be in close proximity to the transmit antenna on the external device. The power link is very sensitive to misalignment between the antennas. For example, some devices use a magnet to manually align them.
Devices based on electromagnetic radiation operate at much higher frequency, 0.5 GHz to 5 GHz. Transmit and receive antennas are on the order of a wavelength. For example, a wavelength is 12.5 cm at 2.4 GHz. Therefore, transmit and receive antennas are usually at least a few cm in diameter which is of similar size to those devices based on inductive coupling. As the transmit antenna is comparable to a wavelength, radiated power dominates. The receive antenna is in the far field of the transmit antenna and captures a very small fraction of the radiated power. That is, most of the transmit power is not delivered to the receiver. The link efficiency is very low. In return, the distance between the transmit antenna and the tissue interface is farther, a few cm to 10's of cm, the depth of the implant inside the body is larger, 1 cm to 2 cm, and the link is insensitive to misalignment between antennas. Prior devices are designed using independent transmit and receive matching networks.
The above two prior approaches have a common disadvantage: they require large receive antennas, 1 cm to a few cm. The paper by Poon et al. titled “Optimal Frequency for Wireless Power transmission over Dispersive Tissue” showed that small receive antenna is feasible. The authors show that the optimal transmission frequency for power delivery over lossy tissue is in the GHz-range for small transmit and small receive antennas (a few mm in diameter.) The optimal frequency for larger transmit antenna (a few cm in diameter) and small receive antenna is in the sub-GHz range. That is, the optimal frequencies are in between 0.5 GHz and 5 GHz. Compared with the frequency used in prior devices based on inductive coupling, the optimal frequency is about 2 orders of magnitude higher. For a fixed receive area, the efficiency can be improved by 30 dB which corresponds to a 10 times increase in the implant depth, from a few mm to a few cm. For a fixed efficiency, the receive area can be reduced by 100 times, from a few cm to a few mm in diameter. When the transmit antenna is close to the tissue interface, the separation between the transmit and the receive antenna approximately equals the implant depth. Inside the body, the wavelength is reduced. For example, a wavelength inside muscle is 1.7 cm at 2.4 GHz. Consequently, the transmit-receive separation is on the order of a wavelength. The device operates neither in the near field nor in the far field. It operates in the mid field. Furthermore, the transmit dimension of a few cm will be comparable to a wavelength.